The invention is related to a passive, ferromagnetic shield for a magnetic resonance imaging magnet. In particular, the invention is related to a magnet having a rectangular, ferromagnetic shield with an open top and an open bottom and an associated plurality of magnetic field inhomogeneity compensating devices for ensuring that a high homogeneity magnetic field is generated in a patient examining region within a bore of the magnet.
It is well known to employ magnetic resonance imaging techniques as a diagnostic tool in the treatment of disease. Magnetic resonance imaging equipment, however, is notoriously expensive, in part because a quasi-static or static highly homogeneous magnetic field must be generated within the examination region occupied by a patient's body. Often a superconducting magnet of the type disclosed in U.S. Pat. No. 4,782,671 for Cooling Apparatus for MRI Magnet and Method of Use and assigned to the instant assignee, is employed to generate the static magnetic field. If the quasi-static magnetic field is not homogeneous to within less than twenty parts per million over the diameter spherical volume (DSV), the field inhomogeneities can prevent an accurate depiction of the internal condition of the patient.
In addition to the requirement that the static magnetic field be highly uniform, it must also be relatively intense. As a consequence, it has been found necessary in most commercial embodiments of magnetic resonance imaging equipment to employ superconducting magnets for generating the magnetic field strengths required to image the requisite detail within the human body. As a result of having these high field strengths, it has been found that it is necessary to engage in relatively elaborate shielding of the magnetic field. The Food and Drug Administration of the United States Government has required that the magnetic resonance imaging equipment be provided with an area of exclusion bounded by a five gauss flux line or surface in order to avoid interference with other hospital diagnostic equipment, as well as with devices such as neurostimulators and cardiac pacemakers.
A number of methods have been employed in the past for shielding magnet resonance imaging magnets. In one method employing active shielding, an additional coil or coils are wound about the superconducting coil for generating a magnetic field which when added to the MRI field substantially reduces the magnetic field external to the imaging unit and thereby reduces the volume bounded by the five gauss surface. The problem with such a system is that it is relatively bulky and expensive to manufacture and often requires the use of extra superconducting magnets. The extra magnets in turn require additional cryogenic refrigeration capacity and the like. This can add as much as thirty percent to the cost of a magnetic resonance imaging system.
Other magnetic resonance imaging systems employ passive magnetic shielding. In one type of passive shielding the magnetic resonance imaging system may be placed within a room having walls consisting of ferromagnetic material which provides return paths for the magnetic flux. The room, of course, cannot be used for anything other than magnetic resonance imaging due to the high flux within the room when the superconducting magnet has current circulating in it. In addition, ferromagnetic rooms are relatively expensive to build and place high structural demands upon the building in which they are situated due to the weight of the ferromagnetic material.
Another approach to limiting the external magnetic field is to use a passive ferromagnetic shield having a plurality of symmetric magnetic return paths exterior to the magnetic resonance imaging magnet as is disclosed in U.S. Pat. No. 4,743,880 to Breneman, et al.
The number of flux return paths may be reduced in order to reduce the cost of the magnet, provide better access to the internal portions of the magnet and provide horizontal plane shielding. Unfortunately, if asymmetric instead of symmetric magnetic flux return paths are employed, it has been found that inhomogeneities are introduced into the magnetic field within the magnet bore which render the magnetic resonance imaging system useless for diagnostic purposes.
In order to provide a high-resolution image using nuclear magnetic resonance equipment, it is important to control precisely the magnitude and direction of the quasi-static magnetic field. The quasi-static magnetic vector field determines in part the frequency at which the hydrogen nuclei precessing within the magnetic field will undergo spin flips evidenced by absorption of radio frequency energy of a pre-selected frequency injected into the examination space. If the field varies in magnitude or in direction, and if a pair of gradient fields are added to it in order to provide spacial localization for the energy absorption signal, the inhomogenieties in the primary quasi-static field will reduce the resolution of the magnetic resonance imaging apparatus to the point at which it is impossible to obtain images of adequate resolution.
Another problem with which users of magnetic resonance imaging magnets are faced is compliance with Food and Drug Administration standards requiring that areas of the hospital, clinic or trailer in which the magnetic resonance magnet is located are not subject to a magnetic field intensity greater than five gauss. As a result, most magnetic resonance imaging superconducting magnets are shielded in order to reduce the volume bounded by the five gauss surface.
The shielding may take the form of a room constructed about the magnet of the type disclosed in U.S. Pat. No. 4,646,046 to Vavrek, et al. for Shielded Room Construction for Containment of Fringe Magnetic Fields. Other shields may be cylinders built about the magnet with closely spaced flux return bars of the type disclosed in U.S. Pat. No. 4,646,045 to Chari, et al. for Aperture Size Disc Shaped End Caps of a Ferromagnetic Shield for Magnetic Resonance Magnets. Still other shielding devices employ multiple flat plates, which provide flux return paths as taught in the octagonal structure disclosed in U. S. Pat. No. 4,590,452 to Ries, et al. for Magnetic Device of Apparatus in Nuclear Spin Tomography With a Shielding Device. Some prior magnets employ ferromagnetic cylindrical shells of the type disclosed in U. S. Pat. No. 4,590,428 to Muller, et al. for Electromagnet for NMR Tomography for shielding.
Other workers in the art have provided shielded magnetic structures wherein the superconducting coil wound therein is not wound on a helix, but rather is wound in a variable fashion in order to compensate for perturbations of the magnetic field by the shield, however, tesseral or off-axis components of the magnetic field cannot be compensated by variations in a substantially helically wound coil. Unfortunately, all of these prior art approaches suffer from one or more drawbacks.
The Burnett, et al. approach in U. S. Pat. No. 4,694,269 for a Magnet System and Method of Its Manufacture requires that the magnet coil be precisely wound in a shape other than a helix so that field perturbations may be compensated for. In some cases, however, customers using magnetic resonance imaging equipment in nonmedical environments may find it unnecessary to provide the type of shielding required by the FDA for use in a medical environment. As a result, if the customer elects to leave the shielding off the magnet in order to reduce cost, the pre-wound corrective coils of Burnett, et al. will introduce perturbations into the internal field in the examination space.
Complete shields of the type disclosed in Muller, et al., U. S. Pat. No. 4,590,428 are difficult to work with, since complete shields are relatively heavy, due to the weight of the ferromagnetic material, such as cold-rolled or hot-rolled steel having a thickness ranging from 1 inch to 21/2 inches. The Muller shield must be removed from the magnet before access can be had to the chambers containing the superconducting coil or the liquid helium or liquid nitrogen The system taught by Ries, et al., U. S. Pat. No. 4,590,452, renders the magnet larger than necessary, which would require that the floor of the building in which the magnet is to be located be reinforced to carry the weight of the magnet. The structure taught in U. S. Pat. No. 4,612,505 to Zijlstra for Nuclear Magnetic Resonance Apparatus employing the extremely long cylindrical bars ranged about the magnet, consumes a great deal of space which would make it undesirable to use the magnet in portable or mobile applications. The magnet of Chari, et al., U.S. Pat. No. 4,646,045 would be relatively expensive to build due to its cylindrically arranged flux return bars. In addition, the Chari magnet itself is completely enclosed, which prevents convenient access to the interior.
What is needed is a magnetic resonance imaging magnet having an easily constructed ferromagnetic return path which is relatively light and a system for compensating for magnetic field inhomogeneities introduced into the examining area of the associated superconducting magnet. The superconducting magnet should generate a solenoidal magnetic field in an examination region which is homogeneous to within less than twenty parts per million to provide a quasi-static field for the production of high resolution images by a magnetic resonance imaging apparatus.